Apparatus and method for probing light absorbing agents in biological tissues

ABSTRACT

A method and apparatus are presented for detecting an effect of interactions of electromagnetic radiation with ultrasound radiation at different locations within a region of interest in a scattering medium to thereby enable imaging of said medium. A plurality of sequences of pulses of ultrasound radiation is transmitted towards a plurality of locations, respectively, in said region of interest within an X-Y plane perpendicular to axes of propagation of the ultrasound pulses. Said region of interest is illuminated with incident electromagnetic radiation of at least one wavelength. The phases of the ultrasound radiation or the phases of the electromagnetic radiation components are appropriately controlled in order to be enable identification of the interactions between the electromagnetic and ultrasound radiation that occur at different locations along the Z-axis. Signals of the electromagnetic radiation, each being frequency modulated by a frequency of the ultrasound radiation, are detected, to thereby enabling the imaging of the region of interest.

FIELD OF THE INVENTION

This invention is generally in the field of non-invasive measurementtechniques, and relates to a process and apparatus for real-time imagingand sensing (probing) light absorbing agents, such as hemoglobin, inbiological tissues.

The following is a list of some prior art patents, documents andarticles which are relevant for the better understanding of thebackground of the invention, as will be described further below:

List of References

1. A. Ishimaru, “Wave Propagation and Scattering in Random Media”, Vol.1, Academic Press (1978)

2. M. Kempe et al., “Acousto-optic tomography with multiply scatteredlight”, J. Opt. Soc. A., 14, 5, 1151 (1997)

3. WO 89/00278

4. U.S. Pat. No. 5,174,298

5. U.S. Pat. No. 5,286,968

6. U.S. Pat. No. 5,212,667

7. U.S. Pat. No. 5,951,481

8. U.S. Pat. No. 6,041,248

9. WO 95/33987

10. Fay A. Marks et al, “Comprehensive approach to breast cancerdetection using light: photon localization by ultrasound modulation andtissue characterization by spectral discrimination”, SPIE, vol. 1888,pp.500-509.

11. G. D. Mahan et al., “Ultrasonic tagging of light: theory”, Proc.Natl. Acad. Sci. USA, 95, 14015, (1998).

12. D. J. Pine et al. “Dynamical correlations of multiply-scatteredlight”, Scattering and Localization of Classical Waves in Random Media,Ping Sheng ed. World Scientific (1990).

13. W. Leutz and G. Maret, “Ultrasonic modulation of multiply scatteredlight”, Physica B, 204, 14-19, (1995).

BACKGROUND OF THE INVENTION

In recent years, much effort has been devoted to find a techniquealternative to Magnetic Resonance Imaging (MRI) or Computed Tomography(CT) for non-invasively probing living biological tissues, such as bodyorgans. MRI and CT involve long procedures and do not always allow realtime analysis of measured data. Low-cost, portable and easy-to-usedevices have been developed based on near infrared spectroscopy of blood(e.g., pulse oximetry). This technique, however, provides only a globalpicture of the tissues with a resolution that does not allow functionalimaging of the tissue and a reliable diagnosis.

It is well-known that hemoglobin can be found in the body in twodifferent oxygenation states—oxyhemoglobin and deoxyhemoglobin—whichhave different light absorption spectra (A. Ishimaru, “Wave Propagationand Scattering in Random Media”, Vol. 1, Academic Press (1978)). In thenear infrared range, (690-900 nm), the absorption coefficients of bothstates of hemoglobin are relatively low. At around 804 mm, both stateshave exactly the same absorption coefficient, and this point is called“the isosbestic point”. Therefore, measurement of blood absorption atthis wavelength gives a direct indication of the blood volume beingtested. At longer wavelengths, the absorption is essentially due tooxyhemoglobin. For example, at or around light wavelengths of 1 micron,the absorption of oxyhemoglobin is more than three times higher thanthat of the deoxyhemoglobin. Hence, absorption at these wavelengths(0.804 μm and 1 μm) gives a direct indication of the ratio between thetwo states of hemoglobin.

Hemoglobin oxygenation provides insight on the proper functioning ofmany body organs such as the brain, breast, liver, heart, etc. Otheragents, such as indocyanin green, present absorption in a definiteregion in the near-infrared range, and can be probed also using infraredlight, deeply inside the tissues.

Light propagating inside a scattering medium has twocomponents—ballistic and diffuse light. The first component does notexperience scattering, while the second corresponds to stronglymulti-scattered light (M. Kempe et al., “Acousto-optic tomography withmultiply scattered light”, J. Opt. Soc. A., 14, 5, 1151 (1997)).Ballistic light intensity decreases with distance in a scattering mediummuch more than that of the diffuse light. Therefore, diffuse light canprovide information on a scattering medium deep inside it.

It is known in the art to use the diffuse (scattered) light to obtaininformation on the optical properties of the medium. This is implementedby utilizing an ultrasound wave focused on the particular region underexamination inside the medium. Generally, this technique consists of thefollowing: If an ultrasound wave propagates through a region in ascattering medium and an electromagnetic wave (such as a laser lightbeam) crosses said region and is strongly diffused thereby, theelectromagnetic wave frequency is shifted by the frequency of theultrasound wave (acousto-optic effect) at the location of said region.In other regions, where no interaction between the light and ultrasoundwaves occurs, the frequency of light is unchanged, and consequently, thedetection of the frequency-shifted electromagnetic wave gives directinformation on the absorption properties of said region.

WO 89/00278 discloses a technique of ultrasound tagging of lightutilizing a continuous ultrasound wave. The manner in which this taggingof light is to be done is, however, physically difficult to implement,since the light detection is obtained using a photo-refractive crystalthat requires extremely high intensities.

The ultrasound tagging of light is disclosed also in the followingpublications: U.S. Pat. Nos. 5,174,298; 5,286,968; 5,212,667; 5,951,481;6,041,248; WO 95/33987; Fay A. Marks et al, “Comprehensive approach tobreast cancer detection using light: photon localization by ultrasoundmodulation and tissue characterization by spectral discrimination”,SPIE, vol. 1888, pp. 500-509; and G. D. Mahan et al., “Ultrasonictagging of light: theory”, Proc. Natl. Acad. Sci. USA, 95, 14015,(1998).

U.S. Pat. No. 5,286,968 discloses a technique of multi-channel analogsignal detection, aimed at obtaining synchronous detection with a CCDcamera. This technique is based on a fast laser modulation.

U.S. Pat. No. 5,212,667 discloses a technique of light imaging in ascattering medium using ultrasound probing and speckle imagedifferencing. According to this technique, coherent laser light impingesonto a scattering medium, disposed between two parallel surfaces, in adirection perpendicular to said surfaces. Light emerging from the mediumis a superposition of a multitude of scattered wavelets, eachrepresenting a specific scattering part. These wavelets are projectedonto the viewing plane of a two-dimensional photodetector array, wherethey interfere with each other, giving rise to a speckle pattern.Ultrasound pulses propagate into the scattering medium in a directionsubstantially parallel to said surfaces, and are focused onto the probedregion, thereby effecting changes in the position of the scatterers andcausing a change in the speckle pattern. This method, however, based asit is on a unidirectional laser beam, has a limited capability ofproviding information on the scattering medium.

U.S. Pat. No. 5,951,481 discloses a technique for non-invasivemeasurement of a substance using ultrasound for modulating light that isback-scattered from the region of interest. Here, pulsed ultrasound anda doublet of light pulses are used, and the detected light is not adiffuse light, but a back-scattered, quasi-ballistic light.

U.S. Pat. No. 6,041,248 discloses a technique for frequency encodedultrasound modulated optical tomography of dense turbid media. Thistechnique utilizes frequency chirped ultrasound and modulatedphotomultiplier.

SUMMARY OF THE INVENTION

There is accordingly a need in the art to facilitate two- orthree-dimensional mapping of a region of interest in a scattering mediumby providing a novel method and apparatus based on the principle ofinteraction of diffused light (light that experienced a large number ofscattering events in a medium) with ultrasound radiation.

The present invention provides for real-time analysis of data indicativeof the detected diffused light affected by said interaction to enablereal-time imaging (less than a few seconds per image) and monitoring ofa region of interest in the medium (e.g., a blood volume), and/or oxygensaturation, as well as other light absorbing agents within the medium.This technique is based on time and spatial multiplexing of light by aplurality of ultrasound waves at differently located sample volumes(points) in a medium, as well as proper fast signal processing.

The main idea of the present invention consists of providing anacousto-optic interaction between electromagnetic waves (e.g., laserlight) and ultrasound pulses in order to localize absorption in a turbidmedium (tissues), and affecting the phase of either one of the light orultrasound signals, or both. By providing a certain phase relationshipbetween ultrasound pulses and/or light signals (the so-called “phasecoding”), the location of interactions along the axis of propagation ofthe ultrasound beam (Z-axis) can be provided. In order to locate theseinteractions in the X-Y plane, the ultrasound beams are directed from aplurality of locations in the X-Y plane. By this, a two orthree-dimensional image of a region of interest can be obtained. Theultrasound pulses used in the technique of the present invention aresinusoidal pulses of several (at least one) cycles. An example of suchpulses is the Doppler mode used in medical ultrasonography. TheseDoppler mode pulses are different from doublet pulses that are typicallyused for echography.

The transmission of ultrasound beams to different locations in the X-Yplane can be implemented by using one or more ultrasound transducers(each operable to periodically transmit ultrasound pulses with a certainphase delay). If a single transducer is used, the X-Y plane is scannedby displacing the transducer. When using a plurality (one- ortwo-dimensional array) of transducers operating in parallel, eachtransducer transmitting ultrasound pulses of a frequency slightlydifferent from that of the other transducers, a power spectrum of thetemporal trace automatically gives the signal of all frequencies. It istherefore possible to translate the signal in the frequency domain intothe transducer's position in the X-Y plane.

Alternatively, a phase-array of ultrasonic transducers, similar to thosetypically used in ultrasonic medical imaging, can be used in order toprovide the spatial frequency and phase coding. To this end, theelectrical signal that is sent to each transducer of the phase-arraycomprises several frequencies, and phase delays are chosen appropriatelyfor each frequency.

In order to allow identification of the interaction between theelectromagnetic and ultrasound radiation components that occur atdifferent locations along the Z-axis, the certain phase relationshipbetween the ultrasound pulses may be obtained by providing differentphases of successive ultrasound pulses. Preferably, in order to obtain asufficient signal-to-noise ratio (SNR) in the detected signal, differentphases of the ultrasound pulses are such that each pulse presents adifferent part of a common sinusoidal signal. Alternatively, ultrasonicpulses with an identical temporal profile may be generated, while thelaser intensity is modulated.

There is thus provided according to one aspect of the present invention,a method of detecting the effect of interactions of electromagneticradiation with ultrasound radiation at different locations within aregion of interest in a scattering medium to thereby enable imaging ofsaid medium, the method comprising the steps of:

(i) generating a plurality of sequences of ultrasound pulses, eachcomprising at least one sinusoidal cycle;

(ii) generating incident electromagnetic radiation of at least onewavelength;

(iii) directing the plurality of sequences of said pulses of ultrasoundradiation towards a plurality of locations, respectively, in said regionof interest within an X-Y plane perpendicular to axes of propagation ofthe ultrasound pulses, while illuminating said region of interest withthe incident electromagnetic radiation, and controlling phases of eitherthe ultrasound radiation pulses or the electromagnetic radiationcomponents, to thereby produce signals of the electromagnetic radiation,each being a frequency modulated by a frequency of the ultrasoundradiation and allowing identification of said interactions that occur atsaid plurality of locations in the X-Y plane and in a plurality oflocation along the Z-axis, and;

(iv) detecting the modulated signals of the electromagnetic radiationand generating data indicative thereof, the analysis of said dataenabling the imaging of the region of interest.

By appropriately analyzing the detected signals, information onabsorbing substances in the region of the medium can be obtained.

According to another aspect of the present invention, there is providedan apparatus for detecting an effect of interactions of electromagneticradiation with ultrasound radiation pulses at different locations withina region of interest in a scattering medium to thereby enable imaging ofsaid medium, the apparatus comprising:

(a) an ultrasound firing unit comprising a transducer arrangementoperable to transmit a plurality of sequences of said pulses ofultrasound radiation to a plurality of locations in said region ofinterest with an X-Y plane perpendicular to the axes of propagation ofthe ultrasound pulses; and an electromagnetic radiation source operableto illuminate said region of interest with incident electromagneticradiation of at least one wavelength, to thereby produce signals of theelectromagnetic radiation, each being a frequency modulated by afrequency of the ultrasound radiation and allowing identification ofsaid interactions that occur at said plurality of locations in the X-Yplane and;

(b) a phase control utility associated either with the ultrasound firingunit or with the electromagnetic radiation source, and operable toaffect either phases of the ultrasound radiation pulses or phases of theelectromagnetic radiation components, to thereby allow identification ofsaid interactions that occur at different locations along the Z-axis;

(c) a detector unit for detecting said modulated signals and generatingdata indicative thereof; and

(d) a control unit for operating said ultrasound firing unit, saidelectromagnetic radiation source, and said phase control utility, thecontrol unit comprising a data processing and analyzing utility foranalyzing the data generated by the detector to enable said imaging.

The technique of the present invention provides for obtaining afunctional image of the region when utilizing the combination ofelectromagnetic and ultrasound radiation. It should be understood thatby means of ultrasound radiation only (i.e., by operating the ultrasoundfiring unit and a suitable detector), a structural image of the regionof interest can be obtained. Hence, by selectively operating theultrasound firing unit and the electromagnetic radiation source withcorresponding detectors, both the functional and structural image can beobtained and registered with each other.

BRIEF DESCRIPTION OF THE DRAWINGS

In order to understand the invention and to see how it may be carriedout in practice, a preferred embodiment will now be described, by way ofnon-limiting example only, with reference to the accompanying drawings,in which:

FIGS. 1A and 1B are schematic illustrations of apparatuses according tothe invention, constructed and operated for affecting the phases of,respectively, ultrasound pulses and light components;

FIG. 2 illustrates an ultrasound pulse sequence (continuous trace)suitable to be used in the present invention, the trace being obtainedby using a DDS (Direct Digital Synthesizer) chip aimed at obtaining theproper signal generation;

FIG. 3 illustrates a sinusoidal trace comprising discontinuities;

FIGS. 4A and 4B illustrate the power spectra corresponding,respectively, to a continuous sinusoidal trace, and to a sinusoidaltrace with phase discontinuities;

FIGS. 5A and 5B illustrate the power spectrum with the noise integrationover the frequency range where the intensity peaks take place for,respectively, the case of the continuous sinusoidal excitation, and thecase of phase discontinuities;

FIGS. 6A to 6D illustrate the scheme of the ultrasound pulse firing andthe logical process of the reshaping procedure, according to theinvention;

FIGS. 7A and 7B illustrate an example of the present invention, wherethe technique of the present invention is applied to a completelyhomogenous Agar sample (without any absorber inside);

FIG. 8 illustrates an experimental power spectrum collected for eachposition in the X-Z plane with the example of FIGS. 7A and 7B;

FIG. 9 illustrates the values of the power spectrum at the ultrasoundfrequency (after background removal) as a function of the position ofthe transducer, presenting a map of the absorption in the example ofFIGS. 7A and 7B;

FIG. 10 illustrates the transversal picture of the photon densityobtained with the technique of the present invention;

FIG. 11 illustrates the experimental results in the form of the powerspectrum at the ultrasound frequency after background removal as afunction of the position of the transducer, for another example of theinvention, where a small absorbing region made of the Agar gel with theaddition of black Agar is introduced;

FIGS. 12 and 13 illustrate yet another experiment, where two transducersoperated in parallel with different frequencies are used;

FIGS. 14A and 14B illustrate, respectively, a structure of a monolithicultrasound probe (suitable to be used in a direct transducers' array)utilizing a single transducer, and the typical profile (pulse envelope)of an ultrasound wave generated by this transducer;

FIGS. 15A and 15B illustrate the principles of operation of a phasearray based structure, wherein FIG. 15A corresponds to the focusing withan array transducer in transmission, and FIG. 15B illustrates theprofile (pulse envelopes) of the phase array transducer operating withthree frequencies;

FIG. 16 illustrates an example of a transducer arrangement according tothe invention aimed at exciting closely located points in the medium byregular size (large) transducers;

FIG. 17 illustrates the light direction towards the region of interestin the medium utilizing optical fibers;

FIG. 18 illustrates the experimental results showing the signal to noiseratio as a function of the integration time for both rough data andfitted data with a Rayleigh function;

FIG. 19 illustrates the laser phase modulation scheme;

FIG. 20 illustrates the time dependence of the laser phase;

FIG. 21 illustrates the use of the continuous phase of the light signal;

FIG. 22 illustrates an example of the phase sequence of the lasermodulation scheme;

FIG. 23 illustrates the simulation results of generating differentsignals corresponding to the different positions of the ultrasoundpulse;

FIG. 24 illustrates the time variations of the detected light signal;and

FIG. 25 illustrates the resulting spectrum showing the peakscorresponding to the modulated signals at different locations along theultrasound propagation axis.

DETAILED DESCRIPTION OF THE INVENTION

Referring to FIG. 1A, there is illustrated an apparatus 10A constructedand operated according to one embodiment of the present invention forultrasound modulated light tomography. The apparatus 10 comprises suchmain constructional parts as an ultrasound firing unit 12 coupled to anoptically turbid medium 13 to be imaged (tissues); an illuminator 14(constituting an electromagnetic radiation source) optically coupled tothe medium 13; a phase control utility 16A, which in the present exampleis associated with the ultrasound firing unit 12; a detector 18; and acontrol unit 20.

The ultrasound firing unit 12 comprises a pulse-ultrasound generator 22(including an electronic beam forming unit, and array of amplifiers),and a transducer arrangement 24. The operation of the ultrasound firingunit 12 is aimed at delivering the proper ultrasound wave within thebody.

A function generator 15 transmits a triggering signal TS to the pulsedultrasound generator 22 and to an analog to digital converter (card) 19.Concurrently, the generator 22 transmits an electrical signal to thetransducer arrangement 24 through the phase control utility 16A tothereby actuate one or more transducers to transmit, respectively, oneor more ultrasound signals 25 into a region of interest in the medium.Each ultrasound signal 25 is transmitted in the form of a sequence ofultrasound pulses in a predetermined manner, as will be described morespecifically further below.

The illuminator 14 comprises one or several laser devices 26 generatingincident radiation of at least one wavelength (preferably in a range of690-900 nm), which propagates towards the region of interest. Laserlight is diffused (scattered) by the medium, and the diffused light 27interacts with the ultrasound signal 25, and the signal resulting fromthis interaction is detected by the detector 18. The electric output ofthe detector 18 is directed to the analog-to-digital converter through aband-pass filter and amplifier 28, to thereby produce a correspondingdigital signal (presenting measured data) received by the control unit20. A data processing and analyzing utility 30 of the control unitapplies a power spectrum operation to the measured data, and identifiesvariations in light intensity at different frequencies to determine thelocation and nature of objects in the turbid medium. Measurement resultsare then presented on a display 32.

FIG. 1B illustrates an apparatus 10B for ultrasound modulated lighttomography constructed and operated according to another embodiment ofthe present invention. To facilitate understanding, the same referencenumbers are used to identify the components that are common in theexamples of FIGS. 1A and 1B. Here, a phase control utility 16B isassociated with the illuminator 14, and is a part of the functiongenerator 15. The principles of the operation of apparatus 10B will bedescribed more specifically further below with reference to FIG. 19.

The interaction between the light wave and the ultrasound results inthat the frequency of light is shifted by the frequency of theultrasound, and the amount of absorbing agent can be determined from thechange in the absorption obtained at the frequency shifted light signal.The light source, the probed region, and the detector do not have to bespecifically aligned with each other, and can have any geometricconfiguration, provided that enough photons reach the detector. Thisallows multiple-source/detector configurations, with the increase in thesignal to noise ratio and better light filling of the tissues.

The interaction is as follows: The light source emits light of frequencyω into the probed region (region of interest). The ultrasound pulses offrequency Ω_(US) are transmitted into the probed region. The currentlocation(s) of the interaction in the X-Y plane is defined by thecurrent location of the transducer(s). The ultrasound modulated lighthaving a shifted frequency ω+Ω_(US), and non-modulated light having thefrequency ω are received by the detector, which mixes them and generatesa signal modulated at the ultrasound frequency. Hereinafter, theexpression “modulated signal” will mean the signal detected by thedetector and representing the intensity of the ultrasound modulatedlight (as a result from said interaction), and the expression“non-modulated signal” will mean the signal detected by the detector andrepresenting the intensity of the light not modulated by the ultrasound(having no interaction with the ultrasound signal). The word “signal”without specification will refer to both the modulated and thenon-modulated signal.

Light propagating through the medium (patient's body) experiences theabsorption throughout the regions in the medium. If the ultrasound waveis located in the region of the medium, and the frequency of the lightis changed, detector(s) outside the medium can detect light that haspassed through the medium and selectively detect theultrasound-modulated light. The ratio between the modulated signal andthe non-modulated signal is mainly determined by the local absorptionchanges. Non-modulated light originating from the probed region isdetected together with the modulated light. This is highly desirable inorder to remove the influence of global changes in the probed region.

The change in the absorption of said ultrasound modulated light in theprobed region, due, for example, to changes in the oxygenation state ofthe hemoglobin, is represented by an analog signal, that can then betransformed to a digital signal, be processed, and, if desired,visualized. The modulated signal is proportional to the amplitude of thelight field passing through the probed region, from which the absorptionis calculated: the changes in the modulated signal reflect changes ofthe intensity of the light passing through the probed region, which, inturn, reflects changes in the absorption in the probed region. Thesignal can be processed using an analog-to-digital card with a samplingcycle, which is high enough to sample effectively the signal at theultrasound frequency, the digitized signal being transferred to a memoryand then processed using power spectrum techniques, which directly givesthe signal at the ultrasound frequency. By averaging the signal atseveral frequencies nearby the ultrasound frequency, the averagebackground noise can be obtained and is then subtracted from the signalat the ultrasound frequency.

As described above, the basic idea is to use acousto-optic interactionbetween laser light and ultrasound in order to localize absorption inthe turbid medium. The signal that is detected is a coherent mixing ofultrasound modulated (USM) and not-modulated light. This means that thelaser coherence length must be long enough. A good number is between 10cm and 3 m (path length of the photons in the tissues). The lasercoherence causes the creation of speckle light, which is acharacteristic of laser light undergoing a diffusive process caused bythe interference between different wavelets originating from the samelaser source. The signal is embedded within the speckle.

It is known that in order to obtain a good signal to noise ratio (SNR),the number of speckles “seen” by the detector should be reduced (M.Kempe et al., “Acousto-optic tomography with multiply scattered light”,J. Opt. Soc. Am. A, 14, 5, 1151, (1997)). The speckle decorrelation timeτ_(speckle) is the typical time at which the speckle is moving, or, moreprecisely, the time at which there is a change in the relative phases ofthe interfering wavelets. If τ_(speckle) were infinite, the modulatedsignal (light signal interacted with the ultrasound signal) would be aninfinitely long sinusoidal signal without any phase and amplitudedisruption. However, if this time τ_(speckle) is finite, which is thepractical case in living tissues, then the modulated signal presentsphase (and amplitude) disruptions every τ_(speckle) on average. Thesignal at the ultrasound frequency increases linearly with the dataacquisition duration τ_(acq) until τ_(acq) is approximately equal toτ_(speckle). Then, this ultrasound signal grows only like the squareroot of the ratio τ_(acq)/τ_(speckle). Since the noise also grows as thesquare root of the signal, it means that, data acquisition durationslonger than τ_(speckle) do not increase the signal to noise ratio. Inpractice, the optimum trace duration τ₀ can be up to two or three timesthe speckle decorrelation time.

As indicated above, the ultrasound signal 25 is a sequence of pulses,where each pulse is a sinusoidal pulse composed of several cycles. Thismeans that a small number of periods (between one and a few tens,depending on the application and frequency) are transmitted, followed bya longer period of silence.

In order to obtain two-dimensional pictures of the region of interest(of the absorbing agent existing therein), it is necessary to locate theultrasound in different zones of the X-Y plane of the probed region.This can be done in the following different ways:

(1) using a single transducer and scanning the X-Y plane in the regionof interest with the ultrasound signal generated by this transducer,

(2) using a one-dimensional array of transducers aligned along theX-axis, operating the transducers to simultaneously transmit ultrasoundsignals of different frequencies, and scanning the X-Y plane with thisarray of transducers; and

(3) using a two-dimensional array of transducers and operating thetransducers to simultaneously transmit ultrasound signals of differentfrequencies.

The scanning can be performed either by displacing the transducers withrespect to the medium, or by using a system of phase-array transducers,where the scanning is performed by electronic means only (namely,changing the direction of the ultrasound wave).

In order to obtain in-depth imaging, the interaction zones should beprovided at different locations along the Z-axis, i.e., the axis ofpropagation of the ultrasound wave. This can be implemented by usingultrasound pulses.

During the pulse to pulse duration, the ultrasound pulse travels acertain distance, and a unique relationship exists between the specificpulse' position inside the medium and the time it spent to reach thisposition. By controlling the delay between the time at which theultrasound transducer emits the ultrasonic pulse and the time at whichthe pulse reaches the region of interest, it is possible to controllongitudinal coordinates (along the Z-axis) of the location of theultrasound pulse. For each delay, there is one specific position in theZ-direction (the probed depth z). The probed depth z is given by thepropagation of the pulse of duration τ_(p) inside the tissues. If v_(s)is the speed of sound in the tissues, then the pulse repetitionfrequency (i.e., the inverse of the time interval between two successivepulses) is given by PRF=v_(s)/z_(max), where z_(max) is the maximumdepth desired. This distance can be divided into n_(z) regions in whichthe ultrasound pulse spends a time τ_(b). If the transducer is operatedsuch that 1/PRF is smaller than the optimum trace duration τ₀, thenthere is no phase disruption in average during a period of several pulsefirings, and the signal at the detector is continuous during severalpulse-to-pulse periods. The effective cumulated time spent by thesuccessive ultrasound pulses in one of the n_(z) regions during the timeτ₀ is τ_(eff)=τ₀/n_(z). The cumulated signal at the detectorcorresponding to a given region is formed of n_(b) blocks of durationτ_(b), wherein n_(b)=τ_(eff)/τ_(b). This corresponds to a signal (afterpower spectrum processing) of bandwidth δf˜n_(z)/τ₀. However, thisbandwidth can be achieved only if the cumulated signal at the detectoris a sinusoid with no phase jump. To this end, the phase control utilityis operated to provide a time delay and a specific phase relationshipbetween successive ultrasound pulses in the sequence.

According to one embodiment of the invention, scanning along the Z-axisutilizes a fast detector: its response time is at least twice thehighest operated ultrasound frequency. In this specific embodiment, thelaser is not modulated and the ultrasound pulse sequence to be generatedby the ultrasound firing unit is programmed in a specific manner, aswill be described below. This technique is aimed at providing maximumSNR and concentrating the signal in the narrowest frequency peak (orpeaks) centered at the ultrasound frequency (or frequencies).

This is done by artificially creating a long time series of collecteddata (detected signals) for each location along the propagation axis(Z-axis). In order to do so, all the data acquired when the successiveultrasound pulses pass through a given location are concatenated. Thistime series of data is called a trace. By doing this, several longsinusoidal traces are obtained at a given frequency, whose powerspectrum shows a narrow peak centered at the ultrasound frequency. Eachsuch trace corresponds to a given location along the Z-axis.

It is important to note that the phase continuity should be kept. Asshown in FIG. 2, this is realized by ensuring that for every series ofultrasonic pulses, each pulse beginning phase is the same as thepreceding pulse ending phase.

The disadvantages of using a sinusoidal trace comprising discontinuitieswill now be described with reference to FIGS. 3, 4A-4B and 5A-5B. FIG. 3illustrates the sinusoidal trace comprising phase discontinuities. Thepower spectrum of such trace obtained by Fourier transformation has apeak centered at the sinusoidal frequency, but with a very large width(and possibly some lateral lobes) and much smaller amplitude (since theenergy is spread over a relatively large frequency range). As a result,the number of frequencies that can be included within a given frequencyrange without cross-talk is much smaller, as compared to the case ofphase continuity.

FIGS. 4A and 4B illustrate the power spectra corresponding,respectively, to the trace in the form of a continuous sinusoid, and tothe trace with phase discontinuity. In the example of FIG. 4A, two peaksat two operating frequencies ω₁ and ω₂ are completely timely separated.In the example of FIG. 4B, the two peaks at the two operatingfrequencies ω₁ and ω₂ are crossed, and the part of the signal (shown indashed lines) is lost due to the cross-talk.

Additionally, since the signal must be integrated over a large frequencyrange, the noise is also integrated over a large frequency range, andtherefore the signal to noise ratio decreases approximately linearlywith the frequency span. FIGS. 5A and 5B illustrate the power spectrumwith the noise integration NI over the frequency range where theintensity peaks take place for, respectively, the case of a continuoussinusoidal excitation (noise integration over a small frequency range),and the case of phase discontinuities (noise integration over a largefrequency range).

Hence, in order to keep the phase continuity by ensuring that for everyseries of ultrasonic pulses, each pulse beginning phase is the same asthe preceding pulse ending phase (FIG. 2), the process is as follows:

The ultrasound pulses are generated at a given repetition ratefrequency, such that each pulse beginning phase is the same as thepreceding pulse ending phase. The ultrasound pulse and the continuouswave laser light propagate within the medium and the interactionsbetween them occurs. The light signals resulting from the interactionsare located outside the medium using one or several fast detectors.Analog data indicative of the detected light is transformed into digitaldata using the analog to digital converter (19 in FIGS. 1A and 1B). Dataindicative of the corresponding location of the ultrasonic pulse(location of interaction) is identified. For each location along theultrasound pulse propagation axis (Z-axis), traces of data are created,by concatenating successive data corresponding to the ultrasonic pulselocation. A power spectrum operation is performed on each of theso-obtained traces, and the background-free peaks' amplitudes arecalculated for each of the different peaks and for each location. Thesedata are further processed and the results are displayed.

The scheme of the ultrasound pulse firing, and the logical process ofthe reshaping procedure (performed by the control unit) will now bedescribed with reference to FIGS. 6A-6D. As shown in FIG. 6A, the pulseis launched into the medium and continuously interacts with photons thatare received by the detector. This interaction occurs for each of then_(z) locations during a period τ_(b). Since the speed of lightpropagation is much higher than that of sound, the light propagationtime is negligible. At each relative time from the ultrasonic pulsefiring, such an interaction occurs at a specific position along theultrasound beam propagation axis (Z-axis). After a time period of 1/PRF,a further (new) pulse P₂ is launched, and so on. The pulse is at thesame location in the medium every 1/PRF time. Since the process iscyclic, it is possible to construct several traces from the acquireddata. Each trace corresponds to a given position along the Z-axis and isobtained by appending the signals of length τ_(b) appearing every 1/PRF.FIG. 6B symbolically represents the signal RS that is received from thedetector. Each individual block from blocks B₁-B₄ corresponds to thesignal coming from the corresponding one of n_(z) regions (n_(z)=4 inthe present example) during the duration τ_(b).

After digitization, the signal is cut into a series of small data seriescorresponding to a duration 1/PRF, and starting at the time when thesuccessive ultrasound pulses are fired. Different rows in FIG. 6Ccorrespond to the different pulse firings. Obviously, the first block ofeach row is related to the same position of the ultrasound pulse insidethe medium. FIG. 6D illustrates the fabrication of the different tracescorresponding to the different locations performed by concatenating thedifferent columns obtained in FIG. 6C.

A critical aspect of this embodiment is that the apparatus is operatedso as to ensure that when the blocks are appended, there is no phasejump. As indicated above with reference to FIG. 2, this is realized byensuring that for every series of ultrasonic pulses, each pulsebeginning phase is the same as the preceding pulse ending phase. Thereare several ways of implementing the same.

One way is to send the waveform directly from the programmable functiongenerator. Another way is to use Direct Digital Synthesis DDStechnology, according to which a sinusoidal waveform is stored in alook-up table (LUT) as a vector of numbers (usually with highprecision). An internal or external TTL clock is sent to the chip. Ateach clock tick, the chip takes a number from the LUT and sends it to aDigital to Analog (D/A) module that transforms the number into a veryprecise voltage.

The clock signal is sent for a period τ_(b) to the DDS, which produces asinusoidal signal during this time period. Then, the clock is stoppedfor a time period of 1/PRF-τ_(b), during which the DDS does not send anysignal to the D/A module. Thereafter, the clock again sends a signal tothe DDS, which starts at the position at which it was stopped. The phaseof the ultrasound is therefore kept as needed.

Construction may be such that each transducer has its own clock.Alternatively, a common clock for all the transducers and one DDS chipper transducer can be used. In this case, each DDS delivers a differentfrequency using the same clock. The advantage of this construction isthat all the DDS are automatically synchronized.

The signal from the DDS is sent to the amplifier that matches thetransducers' needs. The amplifier has a fast enough rise time (typicallyten times shorter than the inverse of the transducer's frequency). Forexample, the amplifier model A078 commercially available from LCFEnterprises Ltd. can be used. The amplified signal is sent to thetransducer.

As indicated above, in order to localize the ultrasound beam in the X-Yplane (parallel to the surface of the transducer arrangement by which itis applied to the medium), a two-dimensional array of an ultrasound beammay be provided, e.g., by using n_(x) time moments of beam transmissionby n_(y) transducers. Each transducer has its own frequency. Eachfrequency corresponds to a position on the X-Y plane. For a totalexcitability bandwidth Δf of the ultrasound transducer, the maximumtotal number of transducers is Δf/δf, wherein δf is the signal peaklinewidth for each frequency. For example, if each ultrasound transduceris of the kind capable of generating pulses with frequencies in therange 1-1.5 MHz, then its bandwidth is Δf=500 kHz. If the signal peaklinewidth δf for each frequency is about 1 kHz, then the maximumpossible number of transducers operating simultaneously, each at adifferent frequency, is 500. Because of the cross-talk effects betweenthe transducers, it is better to choose a smaller number of transducers,e.g., a maximum value of n_(max)=Δf/4δf.

Since a power spectrum operation is performed, all the ultrasoundsignals will appear as individual peaks at their respective frequencies.For example, if 16 transducers are working in parallel, 16 differentpeaks will be present in the power spectrum. Therefore, the processingtime and the signal to noise ratio are not dependent on the number oftransducers. Since the peaks' positions are known, special Fouriertransform techniques can be used, e.g., the so-called zoom Fouriertransforms, chirp Fourier transforms, etc.

In order to improve the signal to noise ratio, it is necessary toaverage the signal obtained for a given position. This cannot be done bysimply adding the temporal traces, but by adding the power spectraobtained after the power spectrum operation.

The entire process consists of the following steps: trace reshaping;power spectrum and background removal; mathematical processing forsignal to noise ratio improvement; reconstruction and visualization. Ifreal-time analysis is required, then moving average can be used orsimilar averaging techniques. By using frequencies where there is noultrasound signal, the white noise background is detected and subtractedfrom the signal.

In order to reduce the constraints on the detectors, and to enable theuse of relatively slow detectors, it is possible to modulate the laserlight intensity. If the laser light is intensity modulated at afrequency Ω_(US)+δ_(ω), wherein Ω_(US) is the ultrasound frequency andδ_(ω) is a frequency shift, the signal at the detector has twocomponents: a component with the frequency Ω_(US)+δω corresponding tothe original (incident) light signal, and a component with the frequencyδ_(ω) corresponding to the modulated signal. If, for example, δ_(ω) ischosen smaller than 40 kHz, then a simple sound card, with 16 or morebits, can be used as the A/D card (19 in FIGS. 1A and 1B). Moreover, ifseveral lasers are used, the δ_(ω)-shift can be chosen different foreach laser, and then one single photodetector or photodetector array canbe used for all the wavelengths. Direct laser modulation is in generaldirectly available with diode lasers using a simple current modulation.In the case of several lasers modulated at frequencies Ω₀+δω_(wli) andseveral transducers emitting ultrasound waves at frequenciesΩ₀+δω_(usj), the signal from the i^(th) wavelength and the j^(th)transducer is located as the detected signal with the frequency(δω_(usj)−δω_(wli)).

Once the signal has been obtained in the frequency domain, the processstarts again, and new traces are processed in a similar way. All thepower spectra are then averaged. The signal to noise ratio grows likethe square root of the number of such power spectra.

As indicated above, the ultrasound firing unit may utilize a directarray of transducers (using the phase control utility), or the knownphase array transducers' arrangement.

Experiments have been conducted to obtain two-dimensional transversaland longitudinal pictures in an Agar gel, tissue-like, phantom. The Agargel (“regular” Agar gel) was prepared by mixing 2.7% in volume ofIntralipid 20% (commercially available from B. Braun Melsungen A G) withwater with unclear Agar Fluka 05040 (possessing an intrinsic scatteringcoefficient of μ_(s)=5 cm⁻¹). The amount of 2 droplets/liter of Pelikan4001 black ink were added in order to increase the absorption. Theoptical and ultrasonic properties of the gel have been matched tobiological tissues.

Different cases have been examined. In the first example, the Agarsample AS was completely homogenous, without any absorber inside. In asecond example (experiment), a small absorbing region was introduced,made of exactly the same Agar gel, but with the addition of carbon black(about 250 mg of carbon black per liter), called “black Agar”. This wasdone by drilling a 9 mm-diameter hole in the Agar block, partly fillingthis hole with the black Agar so as to obtain a 1 cm-height, and thencompletely filling the hole with the regular Agar gel. This absorbingregion was index matched to the surrounding medium both for the lightand the ultrasound beams.

In both examples, the experimental set up shown in FIGS. 7A and 7B andthe following conditions were used: The ultrasound transducer 24 wasV314 model (1.5″ effective diameter) commercially available fromPanametrics. The laser was model S2395 (operating wavelength of 695 nm),commercially available from Melles-Griot. The detector was afiber-coupled photomultiplier R7400U-50, commercially available fromHamamatsu. The analog to digital card was model PCI9812, commerciallyavailable from ADLink. The ultrasound wave was focused at a 1″-distanceunder the sample surface, and the light input and output optical fibersF₁ and F₂ were located 1.8 cm apart. The diameter of the ultrasoundwaist (at 6 dB) was 3 mm. The transducer 24 was physically scanned onthe surface of the sample. A 14×14 matrix of transducer positions waschosen, the grid size being 3 mm. Good coupling was ensured by using anultrasound coupling agent (Medi-pharm). At each point, measured data wasobtained and power spectrum operations were performed and averaged.

As shown in the example of FIGS. 7A and 7B, the scan in the X-Z plane isperformed by physically scanning the ultrasound transducer in theX-direction and by sending pulses of duration 6.4 μs to perform the scanin the Z-direction. In the specific example, the ultrasound frequencywas 1.25 MHz. The scan in the X direction comprised 8 locationsseparated by a 5 mm-distance. The scan in the Z-direction comprised 8locations separated by a 9.6 mm-distance. The sample was made of theregular Agar.

FIG. 8 illustrates a spectrum as collected for each position in the X-Zplane for the first example (completely homogenous Agar sample), andFIG. 9 illustrates the values of the power spectrum at the ultrasoundfrequency (after background removal) as a function of the position ofthe transducer, presenting a map of the absorption. This plot is in theform of a banana-shape that is typical for light transport in adiffusive media. FIG. 10 illustrates the so obtained transversal pictureof the photon density.

FIG. 11 illustrates the experimental results in the form of the map ofabsorption (values of the power spectrum at the ultrasound frequencyafter background removal as a function of the position of thetransducer) corresponding to the second example, where the presence ofthe absorbing agent AB is clearly observed.

As indicated above, in order to get a two-dimensional image of theprobed region (in the X-Y plane), it is possible to operate severaltransducers operating at different frequencies in parallel. In thatcase, the optical signal that is detected carries the information onphotons that passed through each of the localized ultrasound waves. Thefollowing experiment was performed. As shown in FIG. 12, two transducers24A and 24B were operated in parallel with frequencies 1 MHz and 1.001MHz, respectively, at fixed positions. The transducers were coupled to aphantom made of regular Agar gel (index matching gel) that mimicstissues optical and ultrasonic properties. Light was coupled to thesample in a reflection configuration, similarly to the experiment withscanning ultrasound. Other experimental conditions were the same asdescribed above. The power spectrum so obtained (after digitalprocessing) is shown on FIG. 13. Two clear peaks P₁ and P₂ were detectedat the transducers' frequencies. These two peaks represent the signalsof the interaction between light and the localized ultrasound waveswithin the Agar sample.

Hence, by combining the use of several transducers arranged in thespaced-apart in the X-Y plane and operating at different frequencies,with the use of phase controlled ultrasound pulses, it is possible toobtain, in real-time, a three-dimensional picture of the absorbingregions in the tissues. The transducers' array may be of the directarray configuration or phase array configuration. In the case of adirect array, each transducer is independent and provides its own beam.Each transducer has its own, independent control channel. Thisconfiguration is particularly suited for a relatively small number ofchannels. In the case of phase array, all the transducers contribute toall the beams through a coherent effect. This configuration isparticularly suited for a large number of channels. The phase-arrayconcept is well-known in the medical ultrasound community. According tothe present invention, the transducer phase array operates with severalfrequencies, rather than one frequency as conventionally used.

FIG. 14A shows a structure 36A of a monolithic ultrasound probe (directtransducers' array). In the example of FIG. 14A, a single-elementfocused transducer 24A accommodated in a transducer housing 37 is shown.The transducer is operated by electrical signals supplied through anelectrical connector 38. The structure 36A makes use of a piezoelectricelement PE as well as an ultrasonic lens and matching layer LM. Thetypical profile (pulse envelope) P_(P) of an ultrasound wave generatedby the transducer 24A is shown in FIG. 14B.

FIGS. 15A-15B illustrate the principles of operation of a phase arraybased structure 36B. FIG. 15A corresponds to the focusing with an arrayof transducers in transmission. The phase array structure uses anelectronic delay scheme. As shown in FIG. 15A, delay lines DL areapplied to the transducer elements, respectively, to provide delayedelectrical excitations of the transducer elements to thereby generate afocused waveform of a transmitted wave TW at a certain location L in themedium. FIG. 15B illustrates the profile (pulse envelopes P₁, P₂ and P₃)of the phase array probe 24B operating with three different frequenciesand creating several spatially resolved focuses along the Z-axis.

It should be noted that relative delays (or phases) between thedifferent elements of the phase array determine the spatial shape of theenvelope, and not the time variations of the ultrasound wave. In thissense, the time modulation (and therefore the Z-dependence) and thespatial shape of the ultrasound envelope are independent.

Both the monolithic ultrasound probe and the phase array structures areoperable to perform the same function: focusing an ultrasound wave in acertain region in the space. The configurations are equivalent in thatthey are capable of simultaneously focusing different ultrasound pulsesof different frequencies at different locations in the XY plane. Theadvantage of using the phase array is that the focus can beelectronically adjusted, due to the fact that the different delays arenot fixed but are electronically adjustable.

The following is the description of the spatial scanning of the X-Yplane by using the direct transducers' array configuration.

Each transducer delivers an ultrasound pulse of the form:(A  ^(  (2  π  f_(p)t + ϕ_(p))))

wherein ƒ_(p) is the transducer's frequency located at the positionX_(p), Y_(p), 0), A is the ultrasound pulse's amplitude, and Φ_(p) isthe phase of the ultrasound pulse. The pulse each channel delivers is apencil-like beam, with a well-defined diameter (which gives theresolution in the X-Y plane).

The transducer array contains closed-packed transducers transmittingpencil-like beams of a diameter defined by the required resolution.Practically, however, it may be difficult to obtain, since small sizedtransducers typically have a large beam divergence.

FIG. 16 illustrates an example of a transducer arrangement 40 accordingto the invention aimed at solving the above problem, namely, at excitingclosely located points in the medium by regular size (large)transducers. The transducer arrangement comprises transducers' array 42composed of large transducers closed-packed on a plane surface, and tworeflective paraboloids 44A and 44B. The paraboloids are designed suchthat the focus ratio of the paraboloids equals the requested sizereduction and the foci of the two paraboloids coincide.

Thus, a plurality of ultrasound beam generated by the transducers formtogether a tube of a diameter D₁. This tube-like beam impinges onto theconcave surface of the paraboloid 44A and, while being reflectedtherefrom, propagate within a solid angle towards the convex surface ofthe paraboloid 44B, which reflects the beam and shapes it like a tube ofa smaller diameter D₂. The two paraboloids actually act as a reflectivetelescope. The exact orientation of the paraboloid axes depends interalia on the precise size of the transducers' array and the orientationof the exit ultrasound waves. In order to avoid aberration of eachindividual ultrasound beam emitted by each transducer, the paraboloidsmay be replaced by polyhedrons that maps the paraboloid surface so thateach individual beam is reflected by a plane surface. In order to reducethe travelling distance of the ultrasound beams before they exit thetransducer arrangement, the first paraboloid 44A can be removed, and theset of transducers are located on an imaginary paraboloid such that thebeams all meet at the focus of the second paraboloid 44B.

The following is a specific numerical example. Sixteen transducers areused, each having a 1.25″-diameter, a 200 mm curvature. These 16transducers (arranged at a 4×4 matrix), have an operating frequency of 1MHz. This corresponds to a beam diameter of 6.5 mm at 6 dB. Thebeginning of the focal zone starts at about 77 mm from the transducerand finishes at about 191 mm from the transducer. A simple geometricanalysis shows that the use of two paraboloids with focal distances of10 cm and 2 cm, respectively, reduces the cross section of theultrasound beam generated by the 4×4 matrix of transducers from 15 cm toa 5 cm-diameter beam exiting from the transducer arrangement.

The following is the description of the embodiment of the inventionwhere the phase-array transducer configuration is used.

In the phase array configuration, the ultrasound beam delivered by eachtransducer strongly diverges, so that it can interfere with the beamsemitted by the other transducers. When using the conventionalphase-array one-frequency configuration of n transducers, eachtransducer emits a wave of the form:(A  ^(  (2  π  f  t + ϕ_(j))))

wherein f is the transducer's frequency. The total complex amplitude atthe point of coordinates (x,y,z) is:$A^{2}\left\{ {n + {\sum\limits_{\underset{j \neq k}{j,{k = 1}}}^{n}{\exp \quad {\left( {\frac{2\quad \pi \quad {f\left( {d_{j} - d_{k}} \right)}}{v} + \phi_{j} - \phi_{k}} \right)}}}} \right\}$

where d_(j)={square root over ((x−X_(j))²+(y−Y_(j))²+z²)} is thedistance between the point and the j^(th) transducer with coordinatesX_(j), Y_(j), 0), and v is the speed of sound in the tissues.

In order to maximize the intensity at the point z, it is necessary thatall the phase terms equal to zero, i.e.:${{\frac{2\quad \pi \quad {f\left( {d_{j} - d_{k}} \right)}}{v} + \phi_{j} - \phi_{k}} = {0\quad {for}\quad {every}\quad j}},{k.}$

This gives (n−1) independent equations that fix the relative phases ofall the transducers. When dealing with the case of multiple frequenciesaccording to the invention, each transducer emits a wave of the form:$\left( {A\quad {\sum\limits_{p = 1}^{P}^{\quad {({{2\quad \pi \quad f_{p}t} + \phi_{j,p}})}}}} \right)$

wherein P is the total number of frequencies. Hence, the total instantintensity is as follows:$A^{2}\left( {n + {\sum\limits_{p,q,j,k}^{{{\quad {({2\quad {\pi {({f_{p} - f_{q}})}}t})}} + \phi_{j,p} - \phi_{k,q}})}}} \right)$

The oscillating term is not zero in average, only if ƒ_(p)=ƒ_(q). Theaverage intensity at the point (x,y,z) is therefore as follows:$A^{2}\left( {n + {\sum\limits_{p,j,k}^{\quad {({{2\quad \pi \quad {f_{p}{({d_{j} - d_{k}})}}} + \phi_{j,p} - \phi_{k,p}})}}}} \right)$

The average intensity is maximized, when the term under the exponentialis nul, that is:

2πƒ_(p)(d _(j) −d _(k))+Φ_(j,p)−Φ_(k,p)=0

For a specific frequency ƒ_(p), and a given point M_(p) with coordinates(x_(p), y_(p), z_(p)), i.e., a given set of values d_(j), a series of(n−1) equations is obtained that fixes the relative values of the phasesΦ_(j,p). It is thus possible to define, for each frequency, a differentpoint at which the ultrasound intensity will be maximal. The number offrequencies, and consequently the number of points probed in the sample,can be much larger than the number of transducers. The same is valid inthe case of direct transducers' array. In the case of phase-array,however, the waveform in the LUT is more complex, and all the DDS arerelated one to the other. More specifically, the waveform that must beintroduced in the LUT of the j^(th) transducer is not a simplesinusoidal curve, but is as follows:$\left( {A\quad {\sum\limits_{p}^{\quad {({{2\quad \pi \quad f_{p}t} + \phi_{j,p}})}}}} \right)$

wherein the frequencies ƒ_(p) are chosen so that there is minimumcross-talk between the signals after the power spectrum procedure, andthe bandwidth of the transducers is optimally filled. The phases arechosen so as to obtain a uniform array of points in the X-Y plane and toprovide scanning of the Z-direction by this plane. This is obtained bysolving the following equations:

 2πƒ_(p)(d _(j) −d _(k))+Φ_(j,p)−Φ_(k,p)=0

for each position. This calculation is made once, and the resultingwaveforms are then stored in the LUT of each DDS. The rest of the scheme(external clock and phase matching) is identical to the case of thedirect transducers' array.

With regard to array of amplifiers in the phase array case, similarly tothe above-described case of the direct transducers' array, the signalfrom the DDS is sent to the amplifier that matches the transducersneeds. The amplifier must have a fast enough rise time (typically tentimes shorter than the inverse of the transducer's highest frequency).After the amplifier stage, the signal is sent to the transducer.

In the case of the phase array, the transducers' array is atwo-dimensional flat array of identical, small size transducers. Thelarger the number of transducers, the better the resolution. Forexample, for an array made of 16×16 transducers, assuming that eachfrequency needs a bandwidth of 1 kHz in order to avoid cross-talk, anecessary bandwidth for the transducers is about 300 kHz, requiring acavity factor Q of about 3, which is reasonable. The size of eachtransducer is smaller than the ultrasound wavelength. The design andfabrication technology of such arrays is known per se and therefore neednot be specifically described.

With regard to the illuminator, the following should be noted. In orderto obtain the oxygen saturation, it is known to use two or moredifferent wavelengths of incident light. To this end, lasers withseveral constraints, such as a long coherence length and sufficientpower, are used.

Different techniques can be used to couple the light in and out themedium (patient's body) such as direct coupling from the laser (usingmirrors and lenses), the use of fibers, etc., provided trespassing ofthe limit imposed by biological damages is avoided.

If coupling of light from the laser to an optical fiber is used, asshown in FIG. 17, such a fiber F can itself be coupled to a coupler 46with several branches, generally at 48. This provides better filling ofthe tissues with light and reducing the local amount of light deliveredto the tissues while keeping an overall larger amount of light. Aholographic diffuser can be attached at the output of each branch inorder to control the shape of the input light distribution when itenters the tissues.

One of the most important points in any imaging technique is to obtain agood enough signal to noise ratio (SNR). The improvement of the signalto noise ratio in the detected signals according to the inventionconsists of matching the statistical trace length with the speckledecorrelation time τ₀, i.e, the phase-shift is only on average occurringat the time τ₀. By introducing a statistical model of the specklefluctuations, and by fitting the experimental results with the model,the signal to noise ratio can be significantly improved.

According to the statistical model, in order to estimate the signal tonoise ratio at the detector, the signal arriving at the detector shouldbe analyzed. The image at the detector is a speckle pattern. Thispattern fluctuates in time, and its probability distribution iswell-known. The signal also fluctuates between power spectra. Thestatistical model based approach simply averages these spectra. A betteralternative is to use the statistical models and fit the data to themodel.

The advantage of this statistical approach is that it is possible toconverge much faster to the signal value. The procedure consists of thefollowing: A series of n power spectra is obtained. One ultrasoundfrequency is selected, and the values of the n spectra at this frequency(signal) are stored. What is then selected are p points around thisfrequency, where no ultrasound beam is present. By this, a noisereference is defined. The n×p values of the noise (background) arestored. The normalized histograms of the noise (background) and of thesignal are determined, and fitted to the Rayleigh distribution. Then,the noise fitted average is subtracted from the signal fitted average.

Experimental results (taken with an ultrasound probe focused at adistance of 30 mm) are illustrated in FIG. 18, showing two graphs G₁ andG₂ corresponding to the signal to noise ratio as a function of theintegration time for, respectively, the rough data and the fitted datawith a Rayleigh function. The fitted data have a higher S/N than thenon-fitted data, up to a 30% factor (in these data, the background hasbeen removed).

The measurement that is performed by the apparatus of the presentinvention is a measurement related to the local light intensity, and notdirectly to local absorption. In order to retrieve the absorption, it isnecessary to perform a reconstruction procedure. A simple reconstructionprocedure is to simply use the known Beer-Lambert law. However otherreconstruction procedures are possible as well (for example, Monte-Carlosimulation, numerical resolution of Maxwell equations, etc.).

Once the local absorption is retrieved at several wavelengths, thetissue saturation index can be easily retrieved from the ratio of theabsorption coefficients at different wavelengths. This procedure iswell-known in the art and has been extensively used in the field ofnear-infrared spectroscopy.

The apparatus according to the present invention can be used fordiagnostic purposes, namely, for diagnosing an abnormality, ormonitoring a patient in order to detect the appearance of such anabnormality; control of a therapeutic treatment; diagnostic andmonitoring. The device can be used as a monitoring device, for examplein the case of brain trauma, where a secondary injury can appear severalhours after the primary injury. In that case, the device maps the regionsurrounding the trauma and looks for changes in the absorption thatwitness changes in the oxygenation of the tissues. The device can beused for the diagnosis of cancerous tumors. The latter are known togenerate a large number of capillaries that sustain their existence.Therefore, the amount of blood that is present in the area surrounding atumor is much larger than in non-cancerous tissues. Tumors can thereforebe detected and monitored.

The apparatus according to the invention can be used for controllingtherapeutic treatment. If a therapeutic treatment generatesmodifications in the absorption or scattering of tissues locally, eitherusing endogenous or exogenous agents, these changes can be monitoredusing the present invention. For example, local laser ablation orcryogenic treatment of tumors would result in a change in the opticalproperties that can be detected and monitored using the device. Othertherapeutic means such as chemical or radiation therapy can be monitoredin a similar way.

The technique of the present invention can be used to monitortemperature within the body. It has been demonstrated that the Brownianmotion of particles can be very accurately determined using diffusivelight (D. J. Pine et al. “Dynamical correlations of multiply-scatteredlight”, Scattering and Localization of Classical Waves in Random Media,Ping Sheng ed. World Scientific (1990)). Early works on diffusive lightspectroscopy have shown that it was possible to differentiate betweenthe Brownian motion of the particles and the motion occurring due to theultrasound (W. Leutz and G. Maret, “Ultrasonic modulation of multiplyscattered light”, Physica B, 204, 14-19, (1995)).

When there is no or very little particle fluctuations, the opticalsignal obtained at the ultrasound frequency, using the algorithmdescribed above, has a peak whose fill width at half maximum (FWHM) isas narrow as the ultrasound peak FWHM. However, when particles arefluctuating, their velocity due to the local fluctuations will be addedto their velocity due to the ultrasound, and the peak FWHM at theultrasound frequency will increase due to the Doppler effect. Since thelocal temperature is proportional to the square root of the localparticles velocity, this gives a way to locally evaluate thetemperature.

A second effect that also contributes to evaluating the localtemperature is the fact that when particles fluctuate, there is a lossof coherence in the signal translated into an increase in the FWHM ofthe detected signal.

Both effects go in the same direction (increase in the FWHM withincreasing temperature). The precise relation between changes in thefluctuations and ΔT can be obtained empirically (for example, bycomparing these measurements with measurements obtained by invasivemethods). A look-up table (LUT) can then be used for relating thechanges in the FWHM and local temperature. By scanning the ultrasoundfocus, a temperature map can be obtained.

In the above-described examples, the location of the interactionsbetween the light and ultrasound radiation at different points along theZ-axis was achieved by controlling the phase between successiveultrasound pulses. Alternatively, different Z-points can be located bysequentially affecting the phase of the light signal without the need ofcontrolling the phase of the ultrasound pulses, and without the need ofa fast detector.

Turning back to FIG. 1B, the function generator may be operable tomodulate the output intensity of the laser using a phase modulationscheme. In this case, the generation and transmission of ultrasoundpulses may be implemented in the conventional manner. In order to scanthe X-Y plane by the plurality of transducers, the transducers may beoperated either sequentially or with different frequencies of ultrasoundpulses. Additionally, in this case, there is no need for using fastdetector(s), but rather a “slow” detector, such as a CMOS camera, whichcan be used and allows parallel detection of a plurality of lightcomponents (by means of a matrix of pixels). The laser light ismodulated with a fixed frequency and a modulated phase. The fixedfrequency can be chosen arbitrarily, for example near the ultrasoundfrequency so that the difference between both frequencies is smallerthan 100 kHz. The modulated phase is built so that when successiveultrasonic pulses pass through the same position, the laser phase isslightly shifted, the shift being different for each different positionof the ultrasound pulse.

When diffuse light passes through the ultrasound pulse, it gets partlymodulated and the modulated signal time dependence is a sinusoidalsignal, whose argument is the difference between the ultrasound andlaser modulations' arguments. It is therefore composed of a fixfrequency and a modulated phase. An optical detector then detects thisoptical signal.

According to one example, the detector response time is smaller than orequal to the ultrasonic pulse-to-pulse periodicity. Therefore, for agiven ultrasonic pulse firing, the detector receives an optical signalthat is the time integration of several signals corresponding to thedifferent spatial positions of the ultrasound pulse.

For a given ultrasonic pulse position, after a time delay correspondingto the ultrasonic pulse-to-pulse periodicity, a new ultrasound pulse,identical to the previous one, occupies the same spatial position, andso on. Since the laser modulation is built so that there is a slightphase-shift between two successive ultrasonic pulses, thelight-ultrasound interaction signal has a phase that is slightly shiftedby a constant amount at each pulse firing. After several firings, foreach pulse position a sinusoidal modulation is described. By choosingthe laser phase-shift so as to be different for each spatial position ofthe ultrasonic pulse, the period of the sinusoid will be different foreach ultrasound pulse position. The light-ultrasound interaction signaltherefore is composed of a series of sinusoidal signals whosefrequencies are related with a one to one correspondence with the pulseposition, and whose amplitude is proportional to the photon density atthe ultrasound position.

Referring to FIG. 19, there is illustrated the laser phase modulationscheme. Successive ultrasound pulses are represented with the same timeorigin (triggering time). The spatial position of a given ultrasonicpulse is associated with a given time of pulse propagation, namely, thepropagation time divided by the speed of sound in the medium. For agiven position of the ultrasonic pulse, successive pulses experience adifferent laser phase that is shifted by a given amount different foreach position.

FIG. 20 illustrates the time dependence of the laser phase withoutfolding. It is represented by a saw-tooth function with increasingtooth-size, the duration of each such tooth being defined by theultrasound pulse-to-pule periodicity T_(pp).

The phase-shift can be chosen so that the end value of the phase isalways a multiple of 2π. In this case, the phase will be continuous,which presents advantages for the electronic implementation. The laserphase can be chosen so as to be a continuous time function. For example,the laser phase can be chosen to be proportional to t₁t, where t₁ is therelative time from the ultrasound pulse firing trigger, and t is theabsolute time, from the start of the measurement session. In this case,the laser phase will have a quadratic dependence with the relative time.

FIG. 21 illustrates the use of the continuous phase, when both the2π-shift multiple is described during a pulse-to-pulse period and thephase is continuous with the time. The total measurement period MP isdefined as the time necessary to obtain a complete sinusoidal period forthe slowest no-null frequency. The total measurement time will be of theorder of magnitude of one to several measurement periods.

In another example, the detector response time is larger than thepulse-to-pulse time duration T_(pp). In this case, the laser modulationphase may be shifted only every several pulse-to-pulse periods, so thatthe same signal reaches the detector during the period of severalultrasonic pulses. This is an averaging operation.

FIG. 22 illustrates the phase sequence of the laser modulation for thecase of integrating a detector and average signal.

It should be understood that although the use of a single laser isexemplified, several lasers could be used in parallel. In this case, thefixed frequency of each laser can be chosen so that, after powerspectrum, each laser occupies a different spectral band. The laser phasemay be synthesized in real-time, or may be computed in advance andstored in a look-up table for faster access. The laser can be modulatedeither directly in the case of semiconductor lasers, through currentmodulation, or by using an external light modulator (such as anacousto-optics modulator)

It should also be noted that the light phase modulation concept can beutilized with the transducer arrangement composed of a single ultrasonictransducer, several transducers, phase-arrayed or not, operating withdifferent frequencies. If several transducers are used, the fixedfrequency of each transducer can be chosen so that, after powerspectrum, each transducer occupies a different spectral band, each bandcorresponding to a different position in the plane perpendicular to theultrasound wave propagation axis (X-Y plane).

Additionally, the modulation scheme of the laser light can be replacedby that of the detector. This includes the modulation of photomultiplierdynodes, if photomultipliers are used, or a multichannel platemodulation if multichannel plates are used in conjunction with a camera.

The following is the simulation example:

The laser modulation is chosen to be proportional to:

sin(Ω_(US)t+αt₁t+βt)

wherein Ω_(US) is the ultrasound central frequency, t₁ is the relativetime starting from each ultrasonic pulse trigger, t is the time from thestart of the total measurement; α is a positive constant; and β is thedifference frequency between the laser frequency and the ultrasoundfrequency.

The signal after the ultrasound-light interaction is proportional to

sin(φ_(L)(t₁,t))t

with φ_(L)(t₁,t)=αt₁t+βt.

In the present example, the values α and β are chosen of the effectivefrequencies to be: α=3 kHz/s and β=30 kHz, and 6 points along the Z-axis(n=6) are supposed to be obtained. The pulse repetition rate is PRF=30kHz. The averaging is carried out only once for the sake of simplicity.Therefore, the frequency ω₀ is ω₀=PRF=30 kHz. The difference between twoconsecutive frequencies (signal frequencies) is 3 kHz. The total phasevalues are:

Position in the Z axis\Time × 30 μs) 0 1 2 3 4 5 6 7 8 9 10 11 0 0 1 2 34 5 6 7 8 9 10 11 1 0 1.1 2.2 3.3 4.4 5.5 6.6 7.7 8.8 9.9 11 12.1 2 01.2 2.4 3.6 4.8 6 7.2 8.4 9.6 10.8 12 13.2 3 0 1.3 2.6 3.9 5.2 6.5 7.89.1 10.4 11.7 13 14.3 4 0 1.4 2.8 4.2 5.6 7 8.4 9.8 11.2 12.8 14 15.4 50 1.5 3 2.5 6 7.5 9 10.5 12 13.5 15 16.5

Each column corresponds to a new ultrasonic pulse (30 microseconds timedifference between two columns). In the linear form, the same is asfollows:

0,0,0,0,0,0,1,1.1,1.2,1.3,1.4,1.5,2,2.2,2.4,2.6,2.8,3,3,3.3 . . . forone average, where the time difference between two points is 5microseconds.

In the case of two averages, we have:

0,0,0,0,0,0,0,0,0,0,0,0,1,1.1,1.2,1.3,1.4,1.5,1,1.1,1.2,1.3,1.4,1.5,2,2.2,2.4,2.6, . . . .

In the process of simulation, the sinus of the phase is performed andmultiplied by a coefficient that simulates the local absorption(coefficients 1 to 6 for the positions 0 to 5 respectively).

Reference is made to FIG. 23, illustrating the different signalscorresponding to the different positions of the ultrasound pulse. Thesummation of the contributions of all the frequencies gives theintegrated signal received by the detector at each sampling time.

The time difference between the consecutive data is 30 microseconds. Incase of two averages, the time would be 60 microseconds. FIG. 24illustrates the time variations of the signal. This summation has beencalculated for 256 sampling points. This is the slow signal detected bythe detector. It should be understood that, in order to obtain the righttime-scale, the time axis data must be multiplied by 30 microseconds. Apower spectrum is then applied to the data and the resulting spectrum isillustrated in FIG. 25, showing 6 clear peaks, whose magnitude isproportional to the coefficients introduced (1,2,3,4,5,6).

Those skilled in the art will readily appreciate that variousmodifications and changes can be applied to the embodiments of theinvention as herein before exemplified without departing from its scopedefined in and by the appended claims.

What is claimed is:
 1. A method of detecting an effect of interactionsof electromagnetic radiation with ultrasound radiation pulses atdifferent locations within a region of interest in a scattering mediumto thereby enable imaging of said medium, the method comprising: (i)generating a plurality of sequences of ultrasound pulses; (ii)generating incident electromagnetic radiation of at least onewavelength; (iii) transmitting the plurality of sequences of said pulsestowards a plurality of locations, respectively, in said region ofinterest within an X-Y plane perpendicular to axes of propagation of theultrasound pulses, while illuminating said region of interest with theincident electromagnetic radiation, to thereby produce signals of theelectromagnetic radiation, each being a frequency modulated by afrequency of the ultrasound radiation and allowing identifications ofsaid interactions that occur at said plurality of locations in the X-Yplane and in a plurality of location along the Z-axis; (iv) detectingthe modulated signals of the electromagnetic radiation and generatingdata indicative thereof, the analysis of said data enabling the imagingof the region of interest the method being characterized in that: eachof said generated ultrasound pulses comprises at least one sinusoidalcycle and said generated electromagnetic radiation is in the form of amodulated continuous wave, the method thereby enabling control of aphase of either the ultrasound or the electromagnetic radiation toprovide a phase continuity of the detected modulated signals.
 2. Themethod according to claim 1, wherein, in step (iii), said plurality ofsequences are transmitted to the different locations in the X-Y plane byscanning the X-Y plane with a single transducer.
 3. The method accordingto claim 1, wherein, in step (iii), said plurality of sequences aretransmitted to the different locations in the X-Y plane by scanning theX-Y plane with a one-dimensional array of transducers aligned along theX-axis.
 4. The method according to claim 1, wherein, in step (iii), saidplurality of sequences are transmitted to the different locations in theX-Y plane by means of a two-dimensional array of transducers alignedalong the X- and Y-axes.
 5. The method according to claim 1, whereinsaid sequences of pulses have different frequencies.
 6. The methodaccording to claim 3, wherein said array of transducers is amulti-frequency phase array, the method comprising electronicallycontrolling the phases of the ultrasound pulses.
 7. The method accordingto claim 1, comprising controlling the phase of the ultrasound radiationby timely separating pulses of different phases in the sequence ofpulses, the pulses presenting different parts of a common sinusoidalsignal, such that each pulse starts with the ending phase of thepreceding pulse.
 8. The method according to claim 1, comprisingcontrolling the phase of the electromagnetic radiation.
 9. The methodaccording to claim 1, wherein the analysis of the data comprises thesteps of converting data indicative of the analog detected signals intodigital data, reshaping the digital signal to thereby obtain a propertime series of data for each location, and performing a power spectrumoperation on the time series.
 10. The method according to claim 1,wherein the analysis of the data enables determination of information onabsorbing substances in said region of interest.
 11. The methodaccording to claim 1, enabling monitoring of a temperature within themedium.
 12. The method according to claim 11, wherein said monitoringcomprises analyzing the data indicative of the detected modulatedsignals and generating data indicative of temperature mapping of theregion of interest.
 13. The method according to claim 1, wherein theimaging of the region of interest utilizing the interacting ultrasoundand electromagnetic radiations enables obtaining a functional image ofthe region of interest, the method comprising: transmitting ultrasoundradiation to different locations inside the region of interest anddetecting ultrasound radiation returned from said locations, to therebyobtain a structural image of said region of interest registered with thefunctional image.
 14. An apparatus for detecting an effect ofinteractions of electromagnetic radiation with ultrasound radiation atdifferent locations within a region of interest in a scattering mediumto thereby enable imaging of said medium, the apparatus comprising: (i)an ultrasound firing unit comprising a transducer arrangement operableto transmit a plurality of sequence of pulses of ultrasound radiation toa plurality of locations in said region of interest with an X-Y planeperpendicular to the axes of propagation of the ultrasound pulses; andan electromagnetic radiation source operable to illuminate said regionof interest with incident electromagnetic radiation of at least onewavelength, to thereby produce signals of the electromagnetic radiation,each being frequency modulated by a frequency of the ultrasoundradiation; (ii) a phase control utility operable to affect the phase ofthe generated radiation; (iii) a detector unit operable to detect saidmodulated signals and generate data indicative thereof; and (iv) acontrol unit for operating said ultrasound firing unit, saidelectromagnetic radiation source, and said phase control utility, thecontrol unit comprising a data processing and analyzing utility foranalyzing the data generated by the detector to enable said imaging; theapparatus being characterized in that: said ultrasound firing unitoperates to provide each of said pulses of ultrasound radiation in theform of at least one sinusoidal cycle, said electromagnetic radiationsource operates to produce said incident electromagnetic radiation inthe form of a modulated continuous wave, and the control unit isoperable to affect the phase of either the ultrasound radiation or theelectromagnetic radiation to provide a phase continuity of the frequencymodulated signals.
 15. The apparatus according to claim 14, wherein saidtransducer arrangement comprises a single transducer displaceable withinthe X-Y plane.
 16. The apparatus according to claim 14, wherein saidtransducer arrangement comprises a one-dimensional array of transducersaligned in a spaced-apart relationship along the X-axis, and beingdisplaceable along the X-axis.
 17. The apparatus according to claim 14,wherein said transducer arrangement comprises a two-dimensional array oftransducers aligned in a spaced-apart relationship along the X- andY-axes.
 18. The apparatus according to claim 17, wherein said sequencesof ultrasound pulses have different frequencies.
 19. The apparatusaccording to claim 14, wherein said phase control utility is operable tocontrol the phase of the ultrasound radiation such that the sequence ofpulses comprises timely separated pulses of different phases.
 20. Theapparatus according to claim 19, wherein the pulses present differentparts of a common sinusoidal signal, such that each pulse starts with anending phase of a preceding pulse.
 21. The apparatus according to claim16, wherein said transducer arrangement comprises a multi-frequencyphase array.
 22. The apparatus according to claim 16, wherein saidultrasound firing unit comprises a beam shaping unit accommodated in thepath of the ultrasound radiation for reducing a diameter of a spotexcited in the medium by the ultrasound radiation, as compared to adiameter of a beam formed by all ultrasound radiation componentsgenerated by the transducers.
 23. The apparatus according to claim 14,wherein said phase control unit is operable to control the phase of theelectromagnetic radiation.
 24. The apparatus according to claim 14,wherein said control unit comprises an analog to digital converter forprocessing analog output of the detector, and a data processing andanalyzing utility operable to reshape a digital signal indicative of thedetected signals to thereby obtain a proper time series of data for eachlocation.
 25. The apparatus according to claim 14, wherein the controlunit is operable to selectively actuate the ultrasound firing unit, andboth the ultrasound firing unit and the electromagnetic radiationsource, signals detected during the operation of the ultrasound firingunit only being indicative of a structural image of the region ofinterest, and said modulated signals being indicative of a functionalimage of the region of interest.
 26. The apparatus according to claim14, wherein the detector unit comprises a CMOS camera for paralleldetection of a plurality of the electromagnetic radiation components.